Recently, attempts to commercialize photon counting type X-ray computed tomography apparatuses (to be referred to as photon counting X-ray CT apparatuses hereinafter) have been made in the form of the extension of techniques concerning single photon detection (to be referred to as single photon detection techniques hereinafter) in nuclear medicine diagnosis apparatuses such as single photon emission computed tomography apparatuses (to be referred to as SPECT apparatuses hereinafter) and positron emission computed tomography apparatuses (to be referred to as PET apparatuses hereinafter). The single photon detection techniques are roughly classified into two types.
The first single photon detection technique is the following method. First of all, a crystal (scintillator) or the like converts X-rays transmitted through an object into scintillation light. A photodetector such as a photomultiplier tube (to be referred to as a PMT hereinafter) or a silicon photomultiplier (to be referred to as an SiPM hereinafter) detects the scintillation light to extract X-ray photons as electrical signals. The above method is called an indirect conversion type method.
The second single photon detection technique is a method of directly converting X-ray photons transmitted through an object into electrical signals by using a semiconductor detector (this method is also called a direct conversion type method). More specifically, a bias voltage is applied to the two electrodes of the semiconductor detector in advance. When X-ray photons enter the semiconductor detector, electron/hole pairs are generated inside the semiconductor detector. The generated electrons and holes are attracted to the different electrodes, respectively. The electrons reaching the electrode are extracted as electrical signals.
In either of the above methods, since the integral value of the intensity of each extracted electrical signal (to be referred to as a detection signal hereinafter) is proportional to the energy of X-ray photons, detection signals are integrated. Integrating detection signals will calculate the energy of each detected X-ray photon. A difference between a nuclear medicine diagnosis apparatus and a photon counting X-ray CT apparatus is that the flow rate of photons in the photon counting X-ray CT apparatus is much higher than that in the nuclear medicine diagnosis apparatus. In order to reconstruct a medical image by using the photon counting X-ray CT apparatus, it is necessary to perform single photon detection at a rate of, for example, 109 photons/mm2/sec (to be referred to as a count rate hereinafter).
When, however, executing single photon detection with respect to X-ray photons at the above count rate, problems arise concerning the following two types of count losses respectively corresponding to the above two types of single photon detection techniques. The problem in the first single photon detection technique concerns a count loss caused by pileup. Pileup occurs when a plurality of X-ray photons enter a scintillator within the typical decay time (several ns) of scintillation. Pileup is a phenomenon in which a plurality of detection signals respectively corresponding to a plurality of X-ray photons overlap. When pileup occurs, a plurality of X-ray photons are counted as one X-ray photon, resulting in a count loss.
The problem concerning the second single photon detection technique concerns a count loss caused when X-ray photons enter the semiconductor detector during the dead time of the semiconductor detector. The dead time is the time interval from the instant a detection signal is extracted from the semiconductor detector to the instant the semiconductor detector can perform pair generation again. When X-ray photons enter the semiconductor detector in the dead time, since no pair generation occurs, no X-ray photons are counted. Currently, there have been attempts to decrease the number of X-ray photons entering the same semiconductor detector within a unit time by decreasing the size (pixel size) of the semiconductor detector. In such attempts, however, the maximum count rate remains at about 106 photons/mm2/sec.
A reason for such problems concerning the above count loss is that a detection signal with a long decay time constant is integrated to calculate the energy of an X-ray photon entering the X-ray detector. In order to implement a photon counting X-ray CT apparatus, it is necessary to achieve a high count rate. It is however difficult to find the above single photon detection techniques on the extension of techniques concerning nuclear medicine diagnosis apparatuses.
As a method of solving the above problems, for example, a method using an X-ray diffraction phenomenon is available. In this method, a diffraction body is provided on the rear surface side of a collimator. A plurality of X-ray detection elements are provided at predetermined distances from the diffraction body. A polychromatic X-ray entering a collimator is diffracted (scattered) at diffraction angles corresponding to the energies of a plurality of monochromatic X-rays included in the polychromatic X-ray. This diffraction causes the polychromatic X-ray to be diffracted (scattered) in the form of Debye-Scherrer rings on a plurality of X-ray detection elements. A histogram representing the number of X-ray photons with respect to the radii of Debye-Scherrer rings is the overlap of the photon count of diffracted X-rays and the photon count of non-diffracted X-rays (transmitted X-rays). In addition, if the distances between the diffraction body and a plurality of X-ray detection elements are short, since a plurality of Debye-Scherrer rings respectively corresponding to a plurality of monochromatic X-rays overlap each other, it is difficult to obtain the energy spectrum of a polychromatic X-ray.